Thermal ablation is a minimally invasive surgical procedure that has been widely used in treating a variety of inoperable tumors and tumor beds. In thermal ablation, radiative, cryogenic, electrical or acoustic energy is delivered through an applicator to a predetermined location(s) inside a target tumor to induce high temperature that causes irreversible cell injury, leading to apoptosis and eventually coagulative necrosis of the cancer cells. The degree of tissue coagulation or damage depends on both the temperature and the duration of the heating process, as well as the tissue composition.
Thermal therapy creates three different zones of hyperthermic ablation in the target tissue. The first zone (Zone 1), which generally includes the area surrounding the applicator tip or central zone, undergoes coagulative necrosis at temperatures ≥50° C. In Zone 1, protein denaturation, enzyme inactivation, cell membrane rupture, mitochondrial injury, hyperchromasia, and at higher temperatures (e.g. >100° C.), tissue vaporization and carbonization also occur. The area away from the applicator tip with a temperature from 41° C. to 45° C. is called the transitional zone (Zone 2). In Zone 2, the heat-injury to the cells is sub-lethal at such temperatures and is often reversible. The tissue beyond Zone 2 (Zone 3) serves as a heat-sink to the other zones due to the blood flows, and thus has decreased ablation efficacy.
The temperature distribution in the tissue is a function of multiple factors, including the energy delivered, energy-to-heat converting efficiency, thermal conductivity, and perfusion of the tissue. An effective treatment should completely coagulate the whole tumor without tissue charring (i.e., the tumor mass is 100% included in Zone 1), while minimizing damages to the normal tissue on the tumors margins (i.e., maintain the margins in Zone 2). Multiple applicators are often required for an efficient and precise ablation, depending on the size and shape of the tumor.
Common types of energy used in thermal ablation are radiofrequency (RF), microwave, high-intensity focused ultrasound (HIFU), and laser.
Radiofrequency ablation (RFA) has been used in numerous solid organ malignancies and is now a part of standard therapy in several tumors including hepatocellular carcinoma. In RFA, a high frequency alternating electric current is applied to the target tissue using RF electrodes (unipolar or bipolar, single or multi-tined). The varying direction of the electric current causes molecular friction due to ionic agitation resulting in tissue heating up to 100° C. However, for temperature >100° C. the tissue electrical impedance increases, limiting the flow of electric current to the remaining part of the target tissue.
Microwave ablation (MWA) uses antennas to apply electromagnetic waves between 900 to 2500 MHz to the target tissue, which generates heat in the target tissue through the process of dielectric hysteresis (rotating dipoles). MWA techniques are more efficient for ablating tissues with high water content (solid organs and tumors) and temperatures >100° C. can be achieved during ablation.
In HIFU ablation, ultrasound energy is focused to the target tumor to generate temperatures up to 60° C. The acoustic pressure waves cause expansion and relaxation of gaseous nuclei within the cells, leading to the collapse of the cell and nuclear membranes, the mitochondria and the endoplasmic reticulum.
Laser-induced thermal therapy (LITT) generally employs a flexible optical fiber or a diffuser attached to the tip of an optical fiber to deliver high power laser (800-1064 nm) to the target tumor. The tissue temperature increases due to light absorption, dominantly by blood. The size of the thermally induced necrosis formed due to laser coagulation can exceed the light penetration depth since part of the energy that is a function of the temperature gradient diffuses into the surrounding colder tissues. The heat generated due to the local absorption of light depends on the fluence rate (W/cm2), the tissue absorption and scattering properties, as well as the thermal conductivity of the tissue. Since tissue optical properties are wavelength dependent, the ablation volume for a given tumor varies depending on the laser ablation wavelength. While it is challenging to compare the existing thermal ablation techniques, LITT has the unique feature of being able to treat multi-focal diseases while also being compatible with magnetic resonance imaging (MRI) and computed tomography (CT), which are widely used for cancer detection and imaging-guidance during surgeries and thermal ablation procedures. LITT, coupled with imaging (e.g., MRI, CT, and ultrasound) has been successfully demonstrated in a number of studies for treating surgically challenging tumors in the brain, prostate, head and neck, liver, lung, kidney, breast, and bone.
Irrespective of the technique used for tumor ablation, imaging and real-time monitoring of temperature distribution and tissue response are critical for an effective and safe therapy. Image-guidance assists in the precise placement of the treatment probe(s) so that an optimal temperature distribution can be obtained across the whole tumor mass. Imaging (e.g. MRI, CT, and ultrasound) has also been utilized to evaluate the response of tumor to the thermal ablation. Temperature elevation is often monitored during ablation procedures and is used as a feedback to control the thermal dosage. An optimal temperature distribution ensures complete destruction of tumor mass, thus reducing recurrence rate, while avoiding tissue charring and excessive damage to normal tissue and critical structures. Temperatures can be monitored invasively or noninvasively. Invasive methods involve the insertion of thermocouples or fiber optic temperature sensors to provide point measurements. Multiple invasive sensors may be required to obtain a coarse temperature map about the tumor under treatment.
In contrast, noninvasive methods can obtain a temperature distribution within the target tissue without any insertions. Magnetic resonance thermometry (MRT) is one of the few noninvasive options for temperature monitoring during thermal ablation, especially in laser ablation due to its MRI compatibility. MRT is based on T1-relaxation or proton resonance frequency (PRF). Recently, the accuracy and temporal resolution of MRT has been evaluated in vitro. This generally concluded that the speed and accuracy of MRT was sufficient for controlling the thermal dosage for tumor ablation. However, both the T1-weighted and PRF-based MRT are tissue-type dependent. The T1-relaxation is affected by the presence of lipid (fat) molecules, so lipid/fat suppression techniques are often incorporated to minimize such effects. On the other hand, PRF is severely affected by microbubbles (created during ablation) and motion artifacts.
Recently, CT-based thermometry (CTT) has also been investigated for thermal ablation monitoring. The influence of some CT scan parameters on the standard deviation of CT numbers was analyzed. These studies indicated that the standard deviations of CT numbers decreased with an increase in tube current-time product, tube voltage and slice thickness, and with the decrease of collimation thickness. The temperature dependence of these parameters could be used to assess the target tissues at certain temperatures. However, the use of CT-thermometry for ablation monitoring is still in its infancy and may require further in vivo studies for improving the standard deviation of CT numbers in the region of interest. Moreover, CT utilizes ionizing radiation for monitoring which makes it a less attractive technique.
Compared to CTT and MRT, ultrasound monitoring techniques otter the benefits of being portable and less expensive, while also not making use of any ionizing radiation. Tissue heating changes the ultrasound propagation speed and attenuation in the tissue consequently changing the acoustic properties of the tissue. Such heat-induced changes in the target tissue induce time and frequency shifts in the backscattered echo signals. The propagation speed is independent of the tissue coagulation and the attenuation coefficient depends on effects due to tissue coagulation and temperature elevation. Therefore, temperature maps can be reconstructed from the backscattered echo signals. However, ultrasound images are affected by microbubbles that occur in tissues for temperatures >100° C., thus the maximum temperature that can be measured is limited to 100° C.
Among the existing monitoring techniques, MRT is the most commonly used in MRI-guided LITT procedures, especially for the treatment of brain tumors. A temperature accuracy of ±1° C. and a spatial resolution of ±1-2 mm can be achieved with MRT. However, MRI and MRT guided LITT require expensive equipment. The size of the MRI/MRT equipment also limits its use in smaller operation rooms and at remote sites. In addition, MRT and other noninvasive temperature-monitoring techniques use an indirect approach to measure the tissue temperature. These techniques only measure changes in the temperature-dependent tissue properties which are assumed to have changed linearly with temperature. Moreover, MRT, although in some ways superior to both CT and ultrasound, is unable to highlight the destroyed areas during the irradiation procedure directly. The coagulation state of the tissue is usually estimated based on the temperature data. Since different tissue types could coagulate at different temperatures, termination of the treatment based on a critical temperature value alone may lead to incomplete ablation and consequently increase the risk for recurrence.
Therefore, there is a clinical need for a noninvasive or minimally invasive LITT device that can simultaneously ablate the tumor mass and measure the tissue temperature and degree of tissue damage during a laser ablation procedure.